1. Field of the Invention
The present invention is directed to a method, in the form of a pulse sequence, for operating a magnetic resonance imaging tomography apparatus, and in particular to such a pulse sequence which is based on the known echo planar method.
2. Description of the Prior Art
A pulse sequence for operating a magnetic resonance imaging (MRI, or Nuclear Magnetic Resonance, NMR) tomography apparatus is known in the art, referred to as the echo planar method. In this known method, a read-out gradient consisting of a plurality of sub-pulses of alternating polarity and at least one phase coding gradient are activated for each scan after an RF excitation. The phase coding gradient consists of sub-pulses which are activated at each polarity change of the read-out gradient, and the resulting signals are digitized and written into a row of a raw data matrix in the K-space per subpulse of the read-out gradient. The rows of the matrix are ordered according to the phase factors defined by the phase coding gradients. Such a pulse sequence is disclosed in European application 0 076 054, corresponding to U.S. Pat. No. 4,509,015.
FIG. 1 shows a typical, conventional magnetic resonance imaging tomography apparatus and FIGS. 2-7 describe the aforementioned known echo planar method, for the purpose of explaining problems associated with this known method.
The apparatus shown in FIG. 1 includes fundamental field coils 1, 2, 3 and 4 which generate a static magnetic field in which the body of a patient 5 to be examined is disposed, if the device is used for medical diagnostics. Gradient coils are also provided which generate independent, orthogonal magnetic fields in the x, y, and z directions, according to the coordinate axes 6. For clarity, only gradient coils 7 and 8 are shown in FIG. 1, which, in combination with a pair of identical gradient coils disposed opposite thereto (not shown) generate the x-gradient. A set of y-gradient coils (not shown) are disposed parallel to the body 5, above and below the body 5, and a set of z-gradient coils (not shown) are disposed at the head and feet of the patient 5, extending transversely relative to the longitudinal axis of the patient 5.
The apparatus also includes an RF antenna 9, which generates and picks-up the nuclear magnetic resonance signals. The coils 1, 2, 3, 4, 7, 8, and 9 surrounded by the dot-dash line 10 constitute the actual examination instrument. The examination instrument is operated by electrical components which include a power supply 11 for the fundamental field coils 1 through 4, and a gradient fields power supply 12, for supplying power to the gradient coils 7 and 8 as well as the further gradient coils. The RF coil 9 is coupled via a signal amplifier 14, or via a RF transmitter 15, to a process control computer 17. A display 18 is connected to the computer 17 for portraying a visual image of the slice of interest. The components 14 and 15 form a RF unit 16 for signal generation and pick-up. A switch 19 permits switching between the transmission mode and the reception mode.
The principle of image generation according to the known echo planar imaging (EPI) method shall be set forth in greater detail with reference to FIGS. 2 through 7. A more detailed disclosure thereof may be found in the aforementioned European application 0 076 054, and U.S. Pat. No. 4,509,015.
At the beginning of a pulse sequence, an RF excitation pulse RF as shown in FIG. 2 is generated under the influence of a slice selection gradient SS. Nuclear spins in a slice of the examination subject are thus excited. Subsequently, the direction of the gradient SS is inverted, and the negative part of the gradient SS cancels the dephasing of the nuclear spins caused by the positive part of the gradient SS.
After the excitation, a phase coding gradient PC as shown in FIG. 4 is generated in the y-direction, and a read-out gradient RO as shown in FIG. 5 is generated in the x-direction. The read-out gradient RO consists of a preliminary pulse ROV and sub-pulses referenced 0 through 5, having alternating polarity. The sub-pulses of the read-out gradient RO are considered as being square-wave pulses in simplified terms, however, in practice a sine function is usually employed, because this is simpler to achieve in terms of the bandwidths of the components used to generate the pulse.
Due to the changing polarity of the read-out gradient RO, the nuclear spins are alternatingly dephased and rephased, so that a sequence of signals S, shown in FIG. 6, arises. After a single excitation, sufficient signals are thus acquired that the entire Fourier K-space is scanned, i.e., the data obtained in this manner are adequate for the reconstruction of a complete tomogram.
The phase coding gradient PC is briefly activated at each change in the polarity of the read-out gradient RO. The phase relation of the nuclear spins is thus forwarded by one step with each activation. A preliminary phasing gradient PCV is activated before the read-out sequence, the purpose thereof being described below.
The resulting nuclear magnetic resonance signals S are sampled in the time domain, are digitized, and the numerical values acquired in this manner are entered into a raw data matrix. The raw data matrix can be considered to be a measured data space, and since in the exemplary embodiment it is two-dimensional, it constitutes a measured data plane. This measured data space or plane is generally referred to as "K-space" in magnetic resonance imaging tomography.
The information regarding the spatial origin of the signal contributions S, needed for the imaging, is coded in the phase factors, with the relationship between the locus space (i.e., the image) and the K-space being mathematically representable as two-dimensional Fourier transformation. This can be expressed as the following equation: EQU S(k.sub.x,k.sub.y)=.intg..intg..zeta.(x,y)exp(i(k.sub.x x+k.sub.y y))dxdy.
The following definitions apply: ##EQU1## wherein .gamma. is the gyromagnetic ratio, G.sub.x (t') is the momentary value of the read-out gradient RO, and G.sub.y (t') is the momentary value of the phase coding gradient PC.
In the raw data matrix shown in FIG. 7, the row numbers correspond to the numbers of sub-pulses of the read-out gradient shown in FIG. 5. For clarity, only eight rows are shown in FIG. 7; in practice, this number is significantly greater, for example 256.
Due to the step-by-step forwarding of the phase coding gradient PC, the sampling in the K-space ensues in successive rows, beginning with the row 0. The change in polarity of the read-out gradient is taken into consideration in that the measured values are entered in the opposite direction in successive rows, for example, toward the right beginning at the left in row 0, and toward the left beginning at the right in row 1.
An image matrix, on the basis of which an image reconstruction can then ensue, is acquired from the raw data matrix shown in FIG. 7 by two-dimensional Fourier transformation. The Fourier transformation supplies the best results if the measured values allocated to the signal maximum reside on the middle row (i.e., in row 4 in the exemplary embodiment). Image artifacts may otherwise arise. Placing the maximum in the "correct" row is achieved by a pre-phasing of the nuclear spins in the y-direction by the pulse PCV of FIG. 4. This pulse is set so that a rephasing is achieved exactly for the middle row (i.e., row 4 in the exemplary embodiment).
The sequence of the measured Fourier rows in the K-space is thus prescribed in the conventional EPI method. Moreover, the measured Fourier rows must be alternatingly entered into the measurement matrix in positive and negative directions, because the read-out gradient alternates in those directions.
This known technique of data acquisition, however, has two significant disadvantages. First, this known method is susceptible to what are known as "N/2 ghosts" due to the alternating entry into the measurement matrix. These arise even if slight deviations exist from row-to-row. This results, given an image matrix of N.times.N points, in the actual image being again imaged shifted by N/2 points in the positive and negative directions, with respect to the middle of the image matrix, generally the ghost images being of a different intensity than the actual image.
A second disadvantage is that the middle measurement rows in the K-space are only read out in the middle of the read-out sequence, given a symmetrically obtained measurement matrix. These middle rows significantly define the signal-to-noise ratio of the image. The FID envelope which would derive given a free induction decay, and which represents the maximally obtainable amplitude of the individual echoes according to FIG. 6, decays according to the function exp(-t/T.sub.2 *) after the excitation, this function being shown in FIG. 6. The value T.sub.2 * is the time constant for the loss of the phase cohesion of the spins, taking magnetic field inhomogeneities into consideration. At that time at which the middle Fourier rows are read out, the signal amplitude is thus already noticeably diminished in comparison to the beginning of the read-out interval. The signal-to-noise ratio is therefore degraded.